Optically driven therapeutic radiation source having a spiral shaped thermionic cathode

ABSTRACT

A therapeutic radiation source includes a spiral-shaped, laser-heated thermionic cathode. A fiber optic cable directs a beam of radiation, having a power level sufficient to heat at least a portion of the electron-emissive surface to an electron emitting temperature, from a laser source onto the cathode. The cathode generates an electron beam along a beam path by thermionic emission, and strikes a target positioned in its beam path. The target includes radiation emissive material that emits therapeutic radiation in response to incident accelerated electrons from the electron beam. The spiral-shaped conductive element has a plurality of spaced apart turns, and is disposed in a vacuum. An interstitial spacing is defined between adjacent turns, so that heat transfer across the spacing between each adjacent turn is essentially eliminated, thereby substantially reducing heat loss in the cathode caused by thermal conduction.

CROSS-REFERENCE TO RELATED APPLICATIONS

Not Applicable

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH

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REFERENCE TO MICROFICHE APPENDIX

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FIELD OF THE INVENTION

The present invention relates to therapeutic radiation sources, and inparticular to miniaturized, highly efficient, optically-driventherapeutic radiation sources.

BACKGROUND OF THE INVENTION

In the field of medicine, radiation may be used for diagnostic,therapeutic and palliative purposes. For example, the therapeutic use ofradiation such as x-rays and y-rays may involve eradicating malignantcells. Conventional radiation treatment systems used for medicaltreatment, such as linear accelerators that produce high-energy x-rays,utilize a remote radiation source external to the targeted tissue. Abeam of radiation is directed at the target area, for example a tumorinside the body of a patient. The x-rays penetrate the patient's bodytissue and deliver radiation to the cancer cells, usually seated deepinside the body. This type of treatment is referred to as teletherapybecause the radiation source is located at some distance from thetarget. This treatment suffers from the disadvantage that tissuedisposed between the radiation source and the target is exposed toradiation. To reach the cancer cells, the x-rays from an externalradiation source must usually penetrate through normal surroundingtissues. Non-cancerous tissues and organs are therefore also damaged bythe penetrating x-ray radiation.

An alternative treatment system utilizing a point source of radiation isdisclosed in U.S. Pat. No. 5,153,900 issued to Nomikos et al., U.S. Pat.No. 5,369,679 to Sliski et al., U.S. Pat. No. 5,422,926 to Smith et.al., and U.S. Pat. No. 5,428,658 to Oettinger et al., all owned by theassignee of the present application, all of which are herebyincorporated by reference. This system includes a miniaturized,insertable probe capable of producing low power x-ray radiation whilepositioned within or in proximity to a predetermined region to beirradiated. The probe may be fully or partially implanted into, orsurface-mounted onto a desired area within a treatment region of apatient. X-rays are emitted from a nominal, or effective “point” sourcelocated within or adjacent to the desired region to be irradiated, sothat a desired region is irradiated, while irradiation of other regionsare minimized. This type of treatment is referred to as brachytherapy, aword derived from the ancient Greek word for close (“brachy”), becausethe source is located close to or in some cases within the areareceiving treatment.

Brachytherapy offers a significant advantage over teletherapy, becausethe radiation is applied primarily to treat a predefined tissue volume,without significantly affecting the tissue adjacent to the treatedvolume. The term brachytherapy is commonly used to describe the use ofradioactive isotopes which can be placed directly within or adjacent thetarget tissue to be treated. Handling and disposing of suchradioisotopes, however, may impose considerable hazards to both thehandling personnel and the environment. X-ray brachytherapy offers theadvantages of brachytherapy, while avoiding the use of radioisotopes.

X-ray brachytherapy treatment generally involves positioning theinsertable probe into or adjacent to the tumor or the site where thetumor or a portion of the tumor was removed to treat the tissue adjacentthe site with a local boost of radiation. Radiation probes of the typegenerally disclosed in U.S. Pat. No. 5,153,900 typically include ahousing, and a hollow, tubular probe or catheter extending from thehousing along an axis and having a target assembly at its distal end.The probe typically encloses an electron source having a thermioniccathode or a photocathode. The electron source also typically includesan accelerating means for establishing an acceleration potentialdifference between the electron source and the target. The target emitsradiation in response to incident electrons from the electron source.

In conventionally heated thermionic cathodes, a filament is resistivelyheated with a current. This in turn heats the cathode so that electronsare generated by thermionic emission. In a typical conventional x-raymachine, for example, the cathode assembly may consist of a thoriatedtungsten coil approximately 2 mm in diameter and 1 to 2 cm in lengthwhich, when resistively heated with a current of 4 A or higher,thermionically emits electrons. Thermionic cathodes must be stableagainst temperature rise under operation, since they may be subject toseveral thousand degrees centigrade. In a photocathode, a photoemissivesubstance is irradiated by a LED or a laser source. Typically, aflexible fiber optical cable couples light from the LED or laser sourceto the photocathode. The laser beam shining down the fiber optic cableactivates the photocathode which generates free electrons by thephotoelectric effect. Photocathodes may be subjected to several hundreddegrees centigrade.

In order to prevent probe failure, it is important that the electronsource be heated as efficiently as possible, namely that the electronsource reach as high a temperature as possible using as little power aspossible. In conventional x-ray tubes, for example, thermal vaporizationof the cathode filament is frequently responsible for tube failure.Also, the anode heated to a high temperature can cause degradation ofthe radiation output. During relatively long exposures from an x-raysource, e.g. during exposures lasting from about 1 to about 3 seconds,the anode temperature may rise sufficiently to cause it to glowbrightly, accompanied by localized surface melting and pitting whichdegrades the radiation output.

While a photocathode avoids such problems, there are difficultiesinherent in fabricating the photocathode, because photocathodefabrication should preferably be done in a vacuum. A photocathode musthave a sufficient quantum efficiency, where quantum efficiency relatesto the number of electrons generated per incident light quantum. Thedegree of efficiency must be balanced to the intensity of availableincident light. For practical substances, with reasonable quantumefficiencies above 10⁻³, the fabrication of the photocathode should beperformed in a vacuum. U.S. Pat. No. 5,428,658, owned by the assignee ofthe present application and hereby incorporated by reference, disclosesan example of such vacuum fabrication.

It is possible to further increase the efficiency of, and reduce thepower requirements of, miniaturized therapeutic radiation sources asdiscussed above, by using a laser rather than an ohmic current, to heatthe thermionic cathode. U.S. patent application Ser. No. (identified byAttorney Docket Nos. PHLL-155 and hereby incorporated byreference)(hereinafter the “PHLL-155” application) discloses a miniaturetherapeutic radiation source that uses a reduced-power, increasedefficiency electron source. The electron source disclosed in thePHLL-155 application has a laser-heated thermionic cathode,.whichgenerates electrons with minimal heat loss, and which does not require avacuum-fabricated photocathode. The electron source includes athermionic cathode having an electron emissive surface. The PHLL-155application discloses using laser energy to heat the electron emissivesurface of the thermionic cathode, instead of resistively heating theelectron emissive surface of the thermionic cathode. In this way,electrons can be produced in a quantity sufficient to form an electroncurrent necessary for generating therapeutic radiation at the target,while significantly reducing the requisite power requirements for theradiation source.

It is desirable that the surfaces of the thermionic cathodes be heatedto as high a temperature as possible, and as rapidly as possible, i.e.that the surfaces be heated as efficiently as possible. Therefore, oneway of reducing the power requirements for a therapeutic radiationsource, such as the source disclosed in the PHLL-155 application, is tominimize heat loss by the thermionic cathode. Heat loss by laser-heatedthermionic cathodes may include 1) heat lost by thermal conduction; and2) heat loss caused by the portion of incident laser radiation thatremains unabsorbed; and 3) heat loss by thermal radiation. One of thefeatures disclosed in the PHLL-155 application are reflector elements.These reflector elements can reflect back to the thermionic cathodeincident laser radiation that remained unabsorbed by the electronemissive surface of the thermionic cathode, thereby minimizing heat lossdue to unabsorbed incident laser radiation. These reflector elementscannot reduce, however, heat loss that is caused by thermal conductionin the thermionic cathode.

It is an object of this invention to reduce heat loss that is caused bythermal conduction in a laser heated thermionic cathode, thereby furtherincreasing the efficiency of a laser-driven therapeutic radiation sourceand reducing the power requirements therefor. It is another object ofthis invention to provide a thermionic cathode for use in a therapeuticradiation source, where the thermionic cathode is shaped and configuredso as to reduce heat loss caused by thermal conduction within thecathode.

SUMMARY OF THE INVENTION

The invention relates to a highly efficient, miniaturized source oftherapeutic radiation, such as x-rays. The therapeutic radiation sourcehas an optically-driven thermionic cathode that is spiral-shaped. Inthis way, heat loss due to thermal conduction within the thermioniccathode is minimized.

A fiber optic cable directs a beam of radiation, having a power levelsufficient to heat at least a portion of the electron-emissive surfaceto an electron emitting temperature, from a laser source onto a cathode.The cathode generates an electron beam along a beam path by thermionicemission, and strikes a target positioned in its beam path. The targetincludes radiation emissive material that emits therapeutic radiation,such as x-rays, in response to incident accelerated electrons from theelectron beam.

In a preferred embodiment, a substantially rigid housing encloses thethermionic cathode and the target. The housing defines a substantiallyevacuated interior region that extends along the beam path, between aproximal end and a distal end of the housing.

In one embodiment, the spiral-shaped thermionic cathode is made of aspiral-shaped conductive element. The spiral-shaped conductive elementhas a plurality of spaced apart turns, and defines an interstitialspacing between each successive turn of said conductive element. Becausethe spiral-shaped conductive element is enclosed within thesubstantially evacuated interior region, heat transfer across theinterstitial spacing between each adjacent turn of the conductiveelement is essentially eliminated. By minimizing heat lost by thermalconduction, the efficiency of the miniaturized thermionic cathode isincreased.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a schematic block diagram of an overview of one embodiment ofa therapeutic radiation source constructed in accord with the presentinvention.

FIG. 2(a) is an overall, diagrammatic view of one embodiment of atherapeutic radiation source constructed according to the presentinvention.

FIG. 2(b) provides an enlarged view of the radiation generator assembly,and the distal end of the probe assembly, constructed in accordance withthe present invention.

FIG. 3(a) shows a plane view spiral-shaped thermionic cathode,constructed in accordance with the present invention.

FIG. 3(b) shows a side view of a spiral-shaped thermionic cathode,constructed in accordance with the present invention.

DETAILED DESCRIPTION

The present invention is directed to a miniaturized, low powertherapeutic radiation source which includes an electron-beam activatedtherapeutic radiation source, and which uses a laser-heated thermioniccathode. As described in the PHLL-155 application, use of a thermioniccathode that is laser-heated significantly reduces the powerrequirements for such therapeutic radiation sources. The presentinvention features the use of a spiral-shaped thermionic cathode, whichis configured so as to minimize energy lost from the incident laserradiation due to thermal conduction within the thermionic cathode. Inthis way, the power requirements for generating therapeutic radiation insuch miniaturized radiation sources are further reduced.

FIG. 1 is a schematic block diagram of an overview of one embodiment ofa therapeutic radiation source 100, constructed according to the presentinvention, and including a spiral-shaped, laser-heated thermioniccathode. In overview, the system of the present invention includes aradiation generator assembly 102, a source of optical radiation 104, anda probe assembly 106. Preferably, the source of optical radiation 104 isa laser, so that the optical radiation generated by the source issubstantially monochromatic, and coherent. The laser may be a diodelaser, by way of example; however, other lasers known in the art may beused, such as a Nd:YAG laser, a Nd:YVO₄ laser, and a molecular laser.

The radiation generator assembly 102 includes an electron source 108,and a target assembly 110 that includes means for emitting therapeuticradiation in response to incident accelerated electrons from theelectron beam. The electron source 108 includes a spiral-shapedthermionic cathode 109. The probe assembly 106 includes optical deliverystructure 112, such as a fiber optical cable assembly. The opticaldelivery structure 112 directs a beam of laser radiation generated bythe laser 104 onto the electron source 108. The laser beam heats thethermionic cathode 109 in the electron source 108, so as to causethermionic emission of electrons. In a preferred embodiment, thespiral-shaped thermionic cathode has a plurality of spaced apart turns,an interstitial spacing being defined between each successive turn. Heatloss in the cathode due to thermal conduction is minimized, due to thespiral-shaped configuration of the cathode.

Generally, the apparatus of the present invention operates at voltagesin the range of approximately 10 keV to 90 keV, and electron beamcurrents in the range of approximately 1 nA to 100 μA. At thoseoperating voltages and currents, radiation output is relatively low, andthe apparatus may be made small enough to be adapted for implantation inmedical therapeutic applications. In view of the low-level radiationoutput, adequate tissue penetration and cumulative dosage may beattained by positioning the radiation source adjacent to or within theregion to be irradiated. Thus, therapeutic radiation is emitted from awell-defined, small source located within or adjacent to the region tobe irradiated.

FIGS. 2(a) and 2(b) show a diagrammatic view of one embodiment of thetherapeutic radiation source apparatus 200 constructed according to thepresent invention. In the embodiment illustrated in FIG. 2(a), theapparatus 200 includes a laser source 204, a probe assembly 206, and aradiation generator assembly 201. The radiation generator assembly 201includes an electron source 208 that generates an electron beam along abeam path 209, and a target assembly 210 positioned in the beam path. Inthe illustrated embodiment, a high voltage power supply 212 is alsoprovided. The probe assembly 206 couples both the laser source 204 andthe high voltage power supply 212 to the target assembly 210. FIG. 2(a)provides an overall view of the therapeutic radiation source 200,whereas FIG. 2(b) provides an enlarged view of 1) the radiationgenerator assembly 201, and 2) the distal end of the probe assembly 206.

Referring to both FIGS. 2(a) and 2(b), the electron source 208 includesa thermionic cathode 222 having an electron emissive surface. Thethermionic cathode 222 is spiral-shaped, and includes a spiral-shapedconductive element having a plurality of spaced apart turns that definean interstitial spacing between adjacent turns. The conductive elementmay be a wire, by way of example. The conductive element may also be aphotochemically machined flat spiral of cathode material. The spiralarrangement of the wire results in a reduction of conductive heat lossin the cathode.

The electron source 208 also includes means for establishing anaccelerating electric field. In one embodiment, the means forestablishing an accelerating electric field may be the high voltagepower supply 212. The high voltage power supply 212 may establish anacceleration potential difference between the thermionic cathode 222 andthe grounded target element 228, so that electrons emitted from thethermionic cathode 222 are accelerated toward the target element 228,and an electron beam is generated. The electron beam is preferably thin(e.g. 1 mm or less in diameter), and is established along a beam path209 along a nominally straight reference axis that extends to the targetassembly 210. The target assembly 210 is positioned in the beam path209. The distance from the electron source 208 to the target assembly210 is preferably less than 2 mm.

The high voltage power supply 212 preferably satisfies threecriteria: 1) small in size; 2) high efficiency, so as to enable the useof battery power; and 3) independently variable x-ray tube voltage andcurrent, so as to enable the unit to be programmed for specificapplications. Preferably, the power supply 212 includes selectivelyoperable control means, including means for selectively controlling theamplitude of the output voltage and the amplitude of the beam generatorcurrent. A high-frequency, switch-mode power converter can be used tomeet these requirements. The most appropriate topology for generatinglow power and high voltage is a resonant voltage converter working inconjunction with a high voltage, Cockroft-Walton-type multiplier.Low-power dissipation, switch-mode power-supply controller-integratedcircuits (IC) are currently available for controlling such topologieswith few ancillary components. A more detailed description of the powersupply 212 is provided in U.S. Pat. Nos. 5,153,900 and 5,428,658.

The target assembly 210 preferably includes a target element 228 spacedapart from and opposite the electron emissive surface of the thermioniccathode 222, where the target element 228 has at least one radiationemissive element adapted to emit therapeutic radiation in response toincident accelerated electrons from the electron emissive surface of thethermionic cathode 222. In a preferred embodiment, the emittedtherapeutic radiation consist of x-rays, however it should be noted thatthe scope of this invention is not limited to x-rays, and other forms oftherapeutic radiation may also be generated.

In one embodiment, the target element 228 is a small beryllium (Be)window, coated on the side exposed to the incident electron beam with athin film or layer of a high-Z, x-ray emissive element, such as tungsten(W), uranium (U) or gold (Au). By way of example, when the electrons areaccelerated to 30 keV−, a 2 micron thick gold layer absorbssubstantially all of the incident electrons, while transmittingapproximately 95% of any 30 keV−, 88% of any 20 keV−, and 83% of any 10keV−x-rays generated in that layer. In this embodiment, the berylliumtarget element 228 is 0.5 mm thick. With this configuration, 95% of thex-rays generated in directions normal to and toward the target element228, and having passed through the gold layer, are then transmittedthrough the beryllium window and outward at the distal end of the probeassembly 206.

In some forms of the invention, the target element 228 may include amultiple layer film, where the differing layers may have differentemission characteristics. By way of example, the first layer may have anemission versus energy peak at a relatively low energy, and the secondunderlying layer may have an emission versus energy peak at a relativelyhigh energy. With this form of the invention, a low energy electron beammay be used to generate x-rays in the first layer, to achieve a firstradiation characteristic, and high energy electrons may be used topenetrate through to the underlying layer, to achieve a second radiationcharacteristic.

In this embodiment, x-rays can be generated in the target assembly inaccordance with pre-selected beam voltage, current, and target elementcomposition. The generated x-rays pass through the beryllium targetsubstrate with minimized loss in energy. As an alternative to beryllium,the target substrate may be made of carbon, ceramic such as boronnitride, or other suitable material which permits x-rays to pass with aminimum loss of energy. An optimal material for target substrate iscarbon in its diamond form, since that material is an excellent heatconductor. Using these parameters, the resultant x-rays have sufficientenergy to penetrate into soft tissues to a depth of a centimeter ormore, the exact depth dependent upon the x-ray energy distribution.

In another embodiment of the invention, the target may be a solid,high-Z material, with x-rays being emitted in an annular beamperpendicular to the tube axis.

The radiation generator assembly 201, which can be for example 1 to 2cm-in length, extends from the end of the probe assembly 206 andincludes a capsule 230 which encloses the target assembly. According toone embodiment, the radiation generator assembly 201 is rigid in natureand generally cylindrical in shape. In this embodiment the cylindricalcapsule 230 enclosing the radiation generator assembly 201 can beconsidered to provide a substantially rigid housing 230 for the electronsource 208. In one embodiment, the electron source 208 and the targetassembly 210 is disposed within the capsule 230, with the thermioniccathode disposed at an input end of the capsule 230, and the targetassembly 210 disposed at an output end of the housing 230. The capsule230 defines a substantially evacuated interior region extending alongthe beam axis 209, between the thermionic cathode 222 at the input endof the capsule 230 and the target assembly 210 at the output end of thehousing 230. The inner surface of the radiation generator assembly 201is lined with an electrical insulator, or a semiconductor, while theexternal surface of the assembly is electrically conductive. Accordingto a preferred embodiment, the radiation generator assembly 201 ishermetically sealed to the end of the probe assembly, and evacuated.According to another embodiment, the entire probe assembly 206 isevacuated.

The probe assembly 206 couples the laser source 204 and the high voltagepower supply 212 to the target assembly 210. In the illustratedembodiment, the probe assembly 206 includes a flexible, electricallyconductive catheter 205 extending along a probe axis between a proximalend and a distal end of the catheter 205. The probe assembly 206includes optical delivery structure 213 having an originating end 213Aand a terminating end 213B. The terminating end 213B of the opticaldelivery structure 213 is affixed to the radiation generator assembly201.

In a preferred embodiment, the optical delivery structure 213 is aflexible fiber optical cable. In this embodiment, the flexible catheter205 that encloses the fiber optical cable 202 is a small-diameter,flexible, metallic outer tube. In this embodiment, the target assembly210 includes an electrically conductive outer surface. Preferably, boththe metallic tube 205 and the target element 228 are set at groundpotential, in order to reduce the shock hazard of the device. In oneembodiment, the fiber optical cable has a diameter of about 200 microns,and the flexible metallic tube 205 has a diameter of about 1.4 mm.

In a preferred embodiment, the fiber optic cable 213 includes anelectrically conductive outer surface. For example, the outer surface ofthe fiber optic cable 213 may be made conductive by applying anelectrically conductive coating. The electrically conductive outersurface of the fiber optic cable 213 provides a connection to thethermionic cathode 222 from the high voltage power supply 212. In thisembodiment, the radiation generator assembly 201 also has anelectrically conductive outer surface. Preferably, both the flexiblemetallic sheath 205 and the outer conductive surface of the radiationgenerator assembly 201 are set at ground potential, in order to reducethe shock hazard of the device. The flexible sheath 205 couples a groundreturn from the target element 228 to the high voltage power supply 212,thereby establishing a high voltage field between the thermionic cathode222 and the target element 228. In an exemplary embodiment, the fiberoptic cable 213 may have a diameter of about 200 microns, and theflexible metallic sheath 205 may have a diameter of about 1.4 mm. Alayer of dielectric material provides insulation between the outersurface of the fiber optic cable 213 and the inner surface of themetallic sheath 205.

Getters may be positioned within the housing.230. The getters aid increating and maintaining a vacuum condition of high quality. The getterhas an activation temperature, after which it will react with stray gasmolecules in the vacuum. It is desirable that the getter used have anactivation temperature that is not so high that the x-ray device will bedamaged when heated to the activation temperature.

The thermionic cathode 222 has an electron emissive surface, and istypically formed of a metallic material. Suitable metallic materialsforming the cathode 222 may include tungsten, thoriated tungsten, othertungsten alloys, thoriated rhenium, and tantalum. In one embodiment, thecathode 222 may be formed by depositing a layer of electron emissivematerial on a base material, so that an electron emissive surface isformed thereon. By way of example, the base material may be formed fromone or more metallic materials, including but not limited to Group VImetals such as tungsten, and Group II metals such as barium. In oneform, the layer of electron emissive material may be formed frommaterials including, but not limited to, aluminum tungstate and scandiumtungstate. The thermionic cathode 222 may also be an oxide coatedcathode, where a coating-of the mixed oxides of barium and strontrium,by way of example, may be applied to a metallic base, such as nickel ora nickel alloy. The metallic base may be made of other materials,including Group VI metals such as tungsten.

The thermionic cathode is spiral-shaped, configured to minimize heatloss through thermal conduction. For a disc-shaped or planar tungstenthermionic cathode, the percentage of incident radiation that isabsorbed at an incident spot on the cathode is typically about 40%. Ofthe 40% absorbed, however, further losses are caused because of thermalconduction within the cathode. In the present invention, the use of aspiral-shaped cathode is disclosed, wherein heat loss through thermalconduction is minimized, because the cathode is in the shape of a spiralcoil having a plurality of spaced apart turns. Heat loss through thermalconduction is substantially reduced, as compared to heat conductionwithin a disk-shaped thermionic cathode, because no heat transfer occursacross the vacuum between adjacent, spaced apart turns of the conductiveelement forming the cathode.

The fiber optical cable 202 is adapted to transmit laser radiation,generated by the laser source 204 (shown in FIG. 2(a)) and incident onthe originating end of the fiber optical cable assembly, to theterminating end of the fiber optical cable assembly 213. The fiberoptical cable 202 is also adapted to deliver a beam of the transmittedlaser radiation to impinge upon the electron-emissive surface of thethermionic cathode 222. The beam of laser radiation must have a powerlevel sufficient to heat at least a portion of the electron-emissivesurface to an electron emitting temperature so as to cause thermionicemission of electrons from the surface.

In operation, the laser beam shining down the fiber optic cable 213impinges upon the surface of the thermionic cathode 222, and rapidlyheats the surface to an electron emitting temperature, below the meltingpoint of the metallic cathode 222. Upon reaching of the surface of aelectron emitting temperature, electrons are thermionically emitted fromthe surface. The high voltage field between the cathode 222 and thetarget element 228 (shown in FIGS. 3 and 4) accelerates these electrons,thereby forcing them to strike the surface of the target element 228 andproduce x-rays. In one embodiment of the invention, a Nd:YAG laser wascoupled into a SiO2 optical fiber having a diameter of 400 microns. A 20kV power supply was used, and a thermionic cathode made of tungsten wasused. Even with a disc-shaped, planar cathode, only a few watts of powerwas needed to generate over 100 μA of electron current. In anotherexample, an infrared diode laser was used in conjunction with aspiral-shaped, half millimeter etched cathode, to achieve about 100 μAof electron current with only 180 mW of power, thereby substantiallyreducing the power requirements for the apparatus 200.

FIGS. 3(a) and 3(b) illustrate in more detail a spiral-shaped cathode300 constructed in accordance with the present invention. FIG. 3(a)illustrates a planar view of the spiral-shaped cathode 300, whereas FIG.3(b) illustrates a side view. In a preferred embodiment, thespiral-shaped cathode 300 may be fabricated by using photoetchingtechniques known in the art. The spiral-shaped cathode 300 includes aconductive element 310 arranged in a spiral shape. The material formingthe spiral-shaped conductive element is preferably a high melting pointmetal adapted to withstand high temperature uses. Suitable materialsforming the cathode may include tungsten, thoriated tungsten, othertungsten alloys, tantalum, rhenium, thoriated rhenium, and molybdenum.Preferably, the spiral-shaped conductive element 310 forms a planarcoil, although other forms of conductive coils may be used, such ashelical coils. Spiral coils of various shapes can be used. For example,each of the plurality of spaced apart turns may have a substantiallycircular shape, when viewed from the longitudinal direction.Alternatively, the spiral coil may have other transverse sectionalshapes, such as oval, square, or rectangular.

The spiral-shaped conductive element 310 has a plurality of spaced apartturns, which define an interstitial spacing 330 between each successiveturn. The conductive element 310 may have a length of about 2 mm toabout 7 mm, although other dimensions are also within the scope of thisinvention. The distance between adjacent turns of the conductive element310 may be about 25 microns to about 50 microns, although otherdimensions are also within the scope of this invention. Since thespiral-shaped cathode 300 is disposed within the vacuum within thecapsule 230 (shown in FIGS. 2(a) and 2(b)), heat transfer across theinterstitial spacing 330 between adjacent turns of the conductiveelement 310 is essentially eliminated. In this way, heat loss in thethermionic cathode 300 that is caused by thermal conduction issubstantially reduced.

In an exemplary embodiment, the spiral-shaped thermionic cathode 300 wasfabricated using a conductive wire 0.002 mm thick, and 7.4 mm long. Inthis embodiment, the conductive wire defined two spaced-apart turns. Thepower loss caused by thermal conduction was only 0.126 Watts, ascompared to planar, disk-shaped cathodes, in which the power loss due tothermal conduction was about 1.1 Watts. The power loss caused by thermalradiation was about 140 mW.

While the invention has been particularly shown and described withreference to specific preferred embodiments, it should be understood bythose skilled in the art that various changes in form and detail may bemade therein without departing from the spirit and scope of theinvention as defined by the appended claims.

1. A therapeutic radiation source, comprising: A. a radiation generatorassembly, comprising: a. an electron source for emitting electrons togenerate an electron beam along a beam path, said electron sourceincluding a thermionic cathode having an electron emissive surface, andb. a target positioned in said beam path, said target including meansfor emitting therapeutic radiation in response to incident acceleratedelectrons from said electron beam; wherein said thermionic cathodecomprises a spiral-shaped conductive element; B. a source of opticalradiation; and C. optical delivery structure having an originating endand a terminating end and adapted for transmitting to said terminatingend optical radiation generated by said source and incident on saidoriginating end; and wherein said optical delivery structure are adaptedfor directing a beam of said transmitted optical radiation upon asurface of said thermionic cathode; and wherein said beam of opticalradiation has a power level sufficient to heat at least a portion ofsaid surface to an electron emitting temperature so as to causethermionic emission of electrons from said surface.
 2. A therapeuticradiation source according to claim 1, further comprising: asubstantially rigid housing enclosing said thermionic cathode and saidtarget, wherein said housing defines a substantially evacuated interiorregion extending along said beam path between a proximal end and adistal end of said housing.
 3. A therapeutic radiation source accordingto claim 1, wherein said thermionic cathode is disposed at said inputend of said housing.
 4. A therapeutic radiation source according toclaim 1, further comprising a radiation transmissive window at an outputend of said housing, wherein therapeutic radiation emitted from saidtarget is directed through said radiation transmissive window.
 5. Atherapeutic radiation source according to claim 1, wherein saidspiral-shaped conductive element defines a plurality of spaced apartturns.
 6. A therapeutic radiation source according to claim 5, whereinsaid conductive element defines an interstitial space between eachsuccessive turn.
 7. A therapeutic radiation source according to claim 5,wherein said spiral-shaped conductive element forms a planar coil.
 8. Atherapeutic radiation source according to claim 5, wherein saidspiral-shaped conductive element forms a helical coil.
 9. A therapeuticradiation source according to claim 5, wherein the distance betweenadjacent turns of said conductive coil spiral-shaped conductive elementis from about 25 microns to about 50 microns.
 10. A therapeuticradiation source according to claim 5, wherein each of said plurality ofspaced apart turns has a transverse sectional shape that issubstantially circular.
 11. A therapeutic radiation source according toclaim 1, wherein said optical delivery structure comprises a fiberoptical cable.
 12. A therapeutic radiation source according to claim 1,wherein said fiber optical cable has a diameter between about 100microns to about 200 microns.
 13. A therapeutic radiation sourceaccording to claim 5, wherein said spiral-shaped conductive coil elementhas a length between about 2 mm to about 7 mm.
 14. A therapeuticradiation source according to claim 1, wherein the power required forheating said electron emissive surface of said cathode so as to generatean electron beam forming a current of about 2 micro amps is betweenabout 0.1 Watt to about 1.0 Watt.
 15. A therapeutic radiation sourceaccording to claim 1, wherein said optical source is a laser, andwherein said beam of optical radiation is substantially monochromaticand coherent.
 16. A therapeutic radiation source according to claim 1,wherein said therapeutic radiation comprises x-rays.
 17. A therapeuticradiation source according to claim 1, wherein power loss caused bythermal conduction is less than 0.2 Watts.
 18. A therapeutic radiationsource according to claim 17, wherein heat transfer across the spacingbetween each adjacent turn of said conductive element is essentiallyeliminated, thereby substantially reducing in said thermionic cathodeheat loss caused by thermal conduction.
 19. A therapeutic radiationsource according to claim 1, further including means for establishing anaccelerating electric field which acts to accelerate electrons emittedfrom said electron source toward said target.
 20. A therapeuticradiation source according to claim 19, wherein said means forestablishing an accelerating electric field is a power supply.
 21. Atherapeutic radiation source, comprising: A. a radiation generatorassembly, comprising: a. an electron source for emitting electrons togenerate an electron beam along a beam path, said electron sourceincluding a thermionic cathode having an electron emissive surface, andb. a target positioned in said beam path, said target including meansfor emitting therapeutic radiation in response to incident acceleratedelectrons from said electron beam; and c. a substantially rigid housingenclosing said thermionic cathode and said target, wherein said housingdefines a substantially evacuated interior region extending along saidbeam path between an input end and an output end of said housing. B. asource of optical radiation; and C. optical delivery structure having anoriginating end and a terminating end and adapted for transmitting tosaid terminating end optical radiation generated by said source andincident on said originating end, said optical delivery structure beingadapted for directing a beam of said transmitted optical radiation upona surface of said thermionic cathode, wherein said beam of opticalradiation has a power level sufficient to heat at least a portion ofsaid surface to an electron emitting temperature so as to causethermionic emission of electrons from said surface; and wherein saidthermionic cathode comprises a spiral-shaped conductive element having aplurality of spaced apart turns.
 22. A probe having a radiation sourceat a distal end, comprising: A. a probe assembly including an opticaldelivery structure adapted for transmitting optical radiation; B. anoptical source for generating optical radiation directed to an end ofsaid optical delivery structure; C. a radiation source coupled to adistal end of said optical delivery structure, said radiation sourcecomprising a thermionic cathode and a target element; a. wherein thethermionic cathode is responsive to said optical radiation transmittedto said distal end to emit electrons, and wherein said thermioniccathode comprises a spiral-shaped conductive element; and b. whereinsaid target element is responsive to incident electrons emitted fromsaid thermionic cathode to emit radiation; D. means for establishing anaccelerating electric field extending between said electron sourcetoward said target element, the electric field being effective toaccelerate electrons emitted from the thermionic cathode toward saidtarget element; Wherein said optical delivery structure is adapted todirect a beam of optical radiation transmitted therethrough to impingeupon a surface of the thermionic cathode, and wherein said beam oftransmitted optical radiation has a power level sufficient to heat atleast a portion of said surface to an electron emitting temperature soas to cause thermionic emission of electrons from said surface.
 23. Aprobe in accordance with claim 22, wherein said optical source comprisesa laser.
 24. A probe in accordance with claim 22, wherein said radiationsource comprises an x-ray source, and said radiation emitted from saidtarget element comprises x-rays.
 25. A probe in accordance with claim22, wherein said optical delivery structure comprises a fiber opticcable.
 26. A probe in accordance with claim 22, wherein said radiationsource comprises a substantially rigid housing enclosing said thermioniccathode and said target element, wherein said housing defines asubstantially evacuated interior region extending along said beam pathbetween a proximal end and a distal end of said housing.
 27. A probe inaccordance with claim 22, wherein said spiral-shaped conductive elementdefines a plurality of spaced apart turns with an interstitial spacebetween each successive turn.
 28. A probe in accordance with claim 22,wherein said spiral-shaped conductive element forms one of a planar coiland a helical coil.
 29. A radiation source, comprising: A. a probeassembly including an optical delivery structure, said optical deliverystructure being adapted for transmitting optical radiation incident on aproximal end thereof to a distal end thereof; B. an optical source forgenerating a beam of optical radiation directed to said proximal end ofsaid optical delivery structure; C. a radiation generator assemblycoupled to said probe assembly, including: a. an electron source,responsive to optical radiation transmitted to said distal end of saidoptical delivery structure, for emitting electrons, the electron sourceincluding a thermionic cathode having an electron emissive surface;wherein said b. a target element including at least one radiationemissive material adapted to emit radiation in response to incidentaccelerated electrons from said electron source; and D. means forproviding an accelerating voltage between said electron source and saidtarget element so as to establish an accelerating electric field whichacts to accelerate electrons emitted from said electron source towardsaid target element; wherein said optical delivery structured is adaptedfor directing a beam of optical radiation transmitted therethrough toimpinge upon a surface of said thermionic cathode, and wherein said beamof transmitted optical radiation has a power level sufficient to heat atleast a portion of said surface to an electron emitting temperature soas to cause thermionic emission of electrons from said surface.
 30. Aflexible probe having an x-ray tube as a distal end, comprising: A. anoptical source for generating optical radiation, B. a flexible opticalfiber having a proximal end and a distal end, and adapted frotransmitting optical radiation incident on said proximal end to saiddistal end; C. an x-ray tube coupled to a distal end of said opticalfiber, comprising a substantially rigid housing enclosing a thermioniccathode and an x-ray target, a. wherein the thermionic cathode isresponsive to said optical radiation transmitted to said distal end toemit electrons; wherein said thermionic cathode comprises aspiral-shaped conductive element; and b. wherein said x-ray target isresponsive to incident electrons emitted from said thermionic cathode toemit x-rays; D. means for establishing an electric field to accelerateelectrons emitted from the thermionic cathode toward said x-ray target;wherein said optical fiber is adapted to direct a beam of opticalradiation transmitted therethrough to impinge upon a surface of thethermionic cathode, and wherein said beam of transmitted opticalradiation has a power level sufficient to heat at least a portion ofsaid surface to an electron emitting temperature so as to causethermionic emission of electrons from said surface.